Gamma cameras, also referred to as nuclear cameras, radioisotope cameras, scintillation cameras and Anger cameras, are often used to measure gamma radiation emitted by a body under examination. By measuring the energy and the location of the gamma ray emissions, an image representative of the gamma radiation emitted from the body under examination can be created.
Gamma rays are produced by virtue of introducing one or more radionuclidies into a region of interest within a patient. These radionuclidies decay, thereby emitting gamma radiation characterized by photons having one or more characteristic energies. For example, Tc-99 m emits photons having a photopeak located at approximately 140.5 keV.
Nuclear gamma cameras typically include one or more detector heads which receive the gamma radiation emanating from a patient. Each detector head includes a scintillation crystal which converts incident radiation to flashes of light. Internal electronics convert each flash of light into an indication of the location and energy of each received incident radiation event. A collimator situated in front of the scintillation crystal is used to limit the field of view of the radiation detector and defines the detector's overall resolution and sensitivity (or efficiency). Typically, the detector head is housed in a radiation blocking material, such as a lead housing.
Each detector head is connected to a gantry systems which rotates the detector head about a subject to obtain a complete data set. Rotation of a single detector head 360 degrees about a subject produces a complete data set. To reduce imaging time gamma cameras often contain two or more detector heads coupled to a single gantry system. An example of a gamma camera having two detector heads coupled to a gantry system is described in U.S. Pat. No. 5,569,924 assigned to Picker International, Inc and titled "Transformable Dual Head SPECT Camera System".
As mentioned above, the resolution and sensitivity of a detector head is primarily governed by the type of collimator selected. Typically, a collimator is comprised of lead septa which define a plurality of passages through which radiation may pass to the scintillation crystal. The lead septa serve to substantially block radiation incident on the collimator which is unable to pass directly through one of the plurality of passages to the scintillation crystal. In this way, an approximate location of an origination point of a particular gamma ray incident on the scintillation crystal may be determined.
The resolution of a collimator relates to the ability to spatially distinguish between the origination point of gamma rays incident on the scintillation crystal. If for instance, the collimator included septa which defined extremely long and narrow passages, then the collimator would likely be considered to have high resolution. More particularly, in order for a gamma ray to have passed directly through the long and narrow passage it would have had to originate from a fairly narrow field of view and thus the collimator is better able to spatially resolve the origination point of incident gamma rays. Conversely, if the collimator included septa which defined extremely short and wide passages, then the collimator would likely be considered to have a low resolution.
Inversely related to the resolution of a collimator is the collimator's efficiency or sensitivity (i.e. as resolution gets better sensitivity gets worse and vice versa). The sensitivity of a collimator is defined to be the ratio of the number of gamma rays that pass through the collimator to the scintillation crystal compared to the number of gamma rays that are incident to the collimator. For example, if there were 1000 gamma rays incident the front surface of the collimator, but only 1 gamma ray passed through to the scintillation crystal then the efficiency would be 0.10%. The higher the sensitivity, the more counts the scintillation crystal receives over time. The number of counts received by a detector head is important in order to be able to distinguish pertinent anatomy in an image over noise. Thus, by having higher sensitivity, the time it takes to obtain the necessary amount of counts for a given image is reduced.
Depending on the object or organ to be imaged and the preferences of a physician, the type of collimator selected for a gamma camera's detector heads may vary from one application to another. For example, when conducting brain flow studies which may typically be imaged with a low count rate, a collimator providing low sensitivity but high resolution may be selected. Alternatively, when imaging regions in which large amounts of noise may be present in the final image, the physician may opt for a collimator providing high sensitivity but low resolution.
Unfortunately, the inherent tradeoffs between efficiency and resolution necessitates that a physician select the type of collimator to be used which balances these factors prior to commencing the imaging procedure. If a collimator desired is different than the collimator(s) already on the gamma camera then extra time must be spent in removing the existing collimator from each detector head and replacing each with the type collimator selected. Further, as many gamma cameras today have at least two or three detector heads, it is usually necessary that two or three collimators of each type be kept in inventory for possible use. This, of course, adds to overall cost and logistics.
Additionally, once a collimator type is selected by a physician, images created using different collimators are not normally available since such images would require that the patient be re-injected with a radionuclide after having switched the collimators. Thus, it is often not possible for the physician to re-image a region with other collimator types.
Another difficulty associated with nuclear imaging is that some of the gamma rays detected by the detector head will have resulted from a scatter event. For instance, in Compton scatter, prior to reaching the scintillation crystal a gamma ray collides with an electron in an outer shell of an atom. The scattered gamma ray is of a lower energy level than the incident gamma ray and also experiences a change in direction. Detection of the scattered gamma ray by the scintillation crystal will in many cases detract from overall image quality as the incident direction of the gamma ray is no longer truly indicative of its origination point within the patient being imaged.
Techniques to compensate for the detection of Compton scatter have been proposed. For instance, one technique is described in U.S. Pat. No. 5,633,500 ('500) assigned to Picker International, Inc. titled "Split Window Scatter Correction". In the '500 patent, the number of counts of gamma rays received in several different energy windows are tracked to estimate and make correction for the contribution of scatter in the energy window containing image data. Determining a precise amount of scatter at each location in the energy window containing the image data is typically difficult if not impossible since the scattering events are random and space variant. Thus, while the '500 patent does provide a reliable method of correcting for scatter, such corrections are based on estimates. Further, because gamma rays incident on a detector are tracked and stored for multiple energy windows and not just the energy window containing the image data, additional memory and processing is typically needed.
Therefore, what is needed is a method and apparatus which overcomes the shortfalls discussed above and others.